Measurement of tissue perfusion, i.e. the flow of fluid in tissue, is important for the functional assessment of organs in vivo. Although the terms perfusion and flow are sometimes used interchangeably, perfusion as used herein refers to a diffusable exchange between a fluid and a substance. In relation to tissue, perfusion is tissue specific and refers to the exchange of oxygen, water or nutrients from blood to tissue.
A number of techniques have been developed to measure tissue perfusion in vivo and in vitro. For example, the wash-in or wash-out kinetics of exogenously administered tracers have been used to measure flow. With diffusible tracers, this type of measurement can yield tissue perfusion rates.
There are two basic classes of experiments used to determine perfusion with tracers: (a) measurement of the terminal deposition of a tracer and (b) measurement of a freely diffusible tracer either by wash-in/wash-out kinetics or by the determination of steady-state tissue level of the tracer. An example of a terminal deposition tracer is radiolabelled microspheres.
There are numerous diffusible tracers, detectable by a variety of techniques, which have been used to measure perfusion. These include Xenon, detected by radioactivity or CT scans, and .sup.15 O-water and .sup.18 F-fluoromethane, detected by positron-emission tomography. These techniques all rely on the administration of exogenous tracers, and may require arterial blood sampling for quantification.
Recently, there have been a number of applications of magnetic resonance imaging (MRI) techniques to measure tissue perfusion. MRI provides detailed images of the human body with soft tissue contrast not achievable with prior imaging techniques. Due to the versatility of this modality, non-invasive evaluations can be made of tissue anatomy, pathology, metabolism and flow.
Magnetic resonance (MR) is defined as the enhanced absorption of energy occurring when the nuclei of atoms or molecules within an external magnetic field are exposed to radio frequency (RF) energy at a specific frequency, called the Larmor or resonance frequency. Drs. Bloch and Purcell each received the Nobel Prize for investigating and describing in 1946 the phenomenon of MR in solids and liquids. The characteristics of the MR signal arising from a given nucleus were found to depend on a specific molecular environment of that nucleus and such signal dependence proved ideal for qualitative and quantitative chemical analysis. Moreover, the radio frequencies involved in MR are nonionizing and can penetrate the human body.
Although MR suggested enormous clinical potential for in vivo studies, the potential of the method was limited by its inability to provide spatial localization of the MR signal. Lauterbur resolved the localization problem through the use of magnetic field gradients. Since 1977, various MR techniques have been developed for the generation of two and three dimensional data of a human subject.
The production of an MR image can be summarized by the following steps. First, randomly oriented nuclei are aligned by a powerful uniform magnetic field. Second this alignment of magnetization is disrupted by properly tuned RF pulses. These pulses disrupt or perturb the nuclei alignment. As the nuclei recover their alignment by relaxation processes, they produce radio signals proportional to the magnitude of their initial alignment. Contrast between nuclei develops as a result of the different rates at which each nuclei realigns with the magnetic field. Third, the positions of the nuclei are localized by the application of a spatially dependent magnetic field called a gradient. Fourth, the radio signals produced by the realigning nuclei are measured or read out after a predetermined time has elapsed from the initial RF excitation. Fifth, the measured or read out signals are transformed by means of a Fourier Transform into data having a particular position in the image being generated. For a more complete discussion of MRI methods and equipment, see R. R. Edelman et al., Clinical Magnetic Resonance Imaging, W. B. Saunders (U.S.A.) (1990), which work is incorporated herein by reference.
Application of magnetic resonance imaging (MRI) techniques to measure tissue perfusion has in the past involved the determination of wash-in or wash-out kinetics of tracers such as .sup.2 H-water, .sup.19 F-trifluoromethane, and chelated gadolinium contrast agents. These experiments are analogous to those for detecting radiolabeled tracers and require the administration of exogenous agents.
Another class of MR measurements exist which is aimed at measuring volume fractions of endogenous tissue water. These methods offer the advantage of being entirely noninvasive and allow for unlimited serial measurements of blood flow. An example is the intravoxel incoherent motion (IVIM) imaging technique as described in D. Le Bihan et al., MRI of intravoxel incoherent motions: Applications to diffusion and perfusion in neurologic disorders, Radiology Vol. 161, p. 401-407 (1986), incorporated herein by reference. This technique attempts to generate perfusion contrast based on the microscopic diffusion of tissue water rather than on tracer kinetics. However, while IVIM yields interesting contrast in images, the exact relationship of the measured quantity by this technique to tissue perfusion rates is not yet clear.
Consequently, a need exists for methods which permit non-invasive measurement of perfusion without the need for the administration of exogenous agents and which provides significantly greater resolution than those methods presently employed.
The present invention involves an alternative technique for proton magnetic resonance imaging of perfusion rates using a fluid as a diffusible tracer. In the specific case of measuring brain prefusion, described below in Examples 1 and 2, the method involves labelling proton spins of inflowing water in the arterial blood using magnetic resonance. In the examples, continuous saturation or inversion is performed proximal to the tissue or organ of interest. Continuous inversion may be achieved using an adiabatic excitation. For imaging perfusion in the brain, spins are labelled in the neck region.
Techniques for adiabatically inverting nuclear spins in arterial blood were described in M. Sardashti et al, "Spin-Labelling Angiography of the Carotids by Presaturation and Simplified Adiabatic Conversion," Magnetic Resonance Medical, Vol. 15, pages 192-200 (1990) (Sardashti) and W. T. Dixon et al., Projection angiograms of blood labeled by adiobatic fast passage, Magnetic Resonance Medical, Vol. 3, pps. 454-462 (1986) (Dixon) which Sardashti and Dixon articles are incorporated herein by reference. In Sardashti, the authors utilize labelled blood in order to investigate intraluminal arterial abnormalities. Such labelling technique was not used in measuring perfusion according to the present invention.